Ultrasonic diagnostic imaging of response frequency differing from transmit frequency

ABSTRACT

An ultrasonic diagnostic imaging system and methods are described which produces ultrasonic images from harmonic echo components of a transmitted fundamental frequency. Preferably, a programmable digital filter is used to pass harmonic echo components for image processing to the exclusion of fundamental frequency signals. In a preferred embodiment, artifacts are removed by producing decorrelated replicas of the harmonic signals, which are then combined and used for imaging. To produce an image in the presence of depth dependent attenuation of high frequency echo signals, both fundamental and harmonic echo signals are processed and used to produce an image blended from components of both fundamental and harmonic echo signals.

This is a continuation-in-part of U.S. patent application Ser. No.08/723,483, filed Sep. 27, 1996 and entitled "ULTRASONIC DIAGNOSTICIMAGING WITH CONTRAST AGENTS" and claims the benefit of U.S. ProvisionalApplication No. 60/032,771 filed Nov. 26, 1996.

This invention relates to ultrasonic diagnosis and imaging of the bodyand, in particular, to new methods and apparatus for ultrasonicallyimaging with a response frequency which differs from the transmittedfrequency.

Ultrasonic diagnostic imaging systems have been used to image the bodywith the enhancement of ultrasonic contrast agents. Contrast agents aresubstances which are biocompatible and exhibit uniquely chosen acousticproperties which return readily identifiable echo signals in response toinsonification. Contrast agents can have several properties whichenables them to enhance an ultrasonic image. One is the nonlinearcharacteristics of many contrast agents. Agents have been producedwhich, when insonified by an ultrasonic wave at one frequency, willexhibit resonance modes which return energy at other frequencies, inparticular, harmonic frequencies. A harmonic contrast agent, wheninsonified at a fundamental frequency, will return echoes at the second,third, fourth, and higher harmonics of that frequency.

It has been known for some time that tissue and fluids also haveinherent nonlinear properties. Tissue and fluids will, even in theabsence of a contrast agent, develop and return their ownnon-fundamental frequency echo response signals, including signals atharmonics of the fundamental. Muir and Carstensen explored theseproperties of water beginning in 1980, and Starritt et al. looked atthese properties in human calf muscle and excised bovine liver.

While these non-fundamental frequency echo components of tissue andfluids are generally not as great in amplitude as the harmoniccomponents returned by harmonic contrast agents, they do exhibit anumber of characteristics which may be advantageously used in ultrasonicimaging. One of us (M. Averkiou) has done extensive research into theseproperties in work described in his doctoral dissertation. In thisexposition and other research, the present inventors have seen that themain lobe of a harmonic beam is narrower than that of its fundamental,which they have found has implications for clutter reduction whenimaging through narrow orifices such as the ribs. They have seen thatthe sidelobe levels of a harmonic beam are lower than the correspondingsidelobe levels of the fundamental beam, which they have found hasimplications for off-axis clutter reduction. They have also seen thatharmonic returns from the near field are also relatively less thanreturning energy at the fundamental frequency, which they have found hasimplications for near field clutter rejection. As will be seen, theseproperties may be exploited in the methods and constructed embodimentsof the present invention.

In accordance with the principles of the present invention, anultrasonic imaging system and method are provided for imaging tissue andfluids from response frequencies which differ from the transmittedfrequency, in particular echoes returned from the tissue or fluids at aharmonic of a transmitted fundamental frequency. The imaging systemcomprises a means for transmitting an ultrasonic wave at a fundamentalfrequency, means for receiving echoes at a harmonic frequency, and animage processor for producing an ultrasonic image from the harmonicfrequency echoes.

In a preferred embodiment of the present invention the transmitting andreceiving means comprise a single ultrasonic probe. In accordance with afurther aspect of the present invention, the probe utilizes a broadbandultrasonic transducer for both transmission and reception.

In accordance with yet another aspect of the present invention,partially decorrelated components of received harmonic echoes areproduced and utilized to remove artifacts from the harmonic image,providing clearly defined images of tissue boundaries such as that ofthe endocardium. In a preferred embodiment the partially decorrelatedcomponents are produced by processing the harmonic echoes throughdifferent passbands.

The methods of the present invention include the use of harmonic echoesto reduce near-field or multipath clutter in an ultrasonic image, suchas that produced when imaging through a narrow acoustic window such asthe ribs. In accordance with yet a further aspect of the presentinvention, harmonic and fundamental echoes are blended in a common imageto reduce clutter, image at appreciable depths, and overcome the effectsof depth-dependent attenuation.

In the drawings:

FIG. 1 illustrates in block diagram form an ultrasonic diagnosticimaging system constructed in accordance with the principles of thepresent invention;

FIGS. 2, 3, 4, and 5 illustrate certain properties of harmonic echoeswhich may be advantageously applied to ultrasonic imaging applications;and

FIGS. 6 and 7 illustrate passband characteristics used to explain theperformance of the embodiment of FIG. 1;

FIG. 8 illustrates typical fundamental and harmonic frequency passbandsof an embodiment of the present invention;

FIG. 9 illustrates an FIR filter structure suitable for use in theembodiment of FIG. 1;

FIG. 10 illustrates in block diagram form a portion of a preferredembodiment of the present invention;

FIG. 11 illustrates the operation of the normalization stages of theembodiment of FIG. 10;

FIG. 12 is a block diagram of one of the multiplier accumulators used inthe filters of the embodiment of FIG. 10;

FIG. 13 illustrates typical fundamental and harmonic frequency passbandsof the embodiment of FIG. 10;

FIG. 14 illustrates the blending of fundamental and harmonic signalcomponents into one ultrasonic image; and

FIG. 15 illustrates the passbands of a time varying filter used in theformation of blended images.

Referring first to FIG. 1, an ultrasonic diagnostic imaging systemconstructed in accordance with the principles of the present inventionis shown in block diagram form. A central controller 120 commands atransmit frequency control 117 to transmit a desired transmit frequencyband. The parameters of the transmit frequency band, f_(tr), are coupledto the transmit frequency control 117, which causes the transducer 112of ultrasonic probe 110 to transmit ultrasonic waves in the fundamentalfrequency band. In a constructed embodiment a band of frequencieslocated about a central frequency of 1.67 MHz is transmitted. This islower than conventional transmitted imaging frequencies, which generallyrange from 2.5 MHz and above. However, use of a typical transmitfrequency of 3 or 5 MHz will produce harmonics at 6 and 10 MHz. Sincehigher frequencies are more greatly attenuated by passage through thebody than lower frequencies, these higher frequency harmonics willexperience significant attenuation as they return to the probe. Thisreduces the depth of penetration and image quality at greater imagingdepths, although the harmonic signals, created as they are during thepropagation of the transmitted wave through tissue, do not experiencethe attenuation of a full round trip from the transducer as thefundamental signals do. To overcome this problem, the central transmitfrequency in the illustrated embodiment is below 5 MHz, and preferablybelow 2.5 MHz, thereby producing lower frequency harmonics that are lesssusceptible to depth dependent attenuation and enabling harmonic imagingat greater depths. A transmitted fundamental frequency of 1.67 MHz willproduce second harmonic return signals at 3.34 MHz in the illustratedembodiment. It will be understood, of course, that any ultrasonicfrequency may be used, with due consideration of the desired depth ofpenetration and the sensitivity of the transducer and ultrasound system.

The array transducer 112 of the probe 110 transmits ultrasonic energyand receives echoes returned in response to this transmission. Theresponse characteristic of the transducer can exhibit two passbands, onearound the fundamental transmit frequency and another about a harmonicfrequency in the received passband. For harmonic imaging, a broadbandtransducer having a passband encompassing both the transmittedfundamental and received harmonic passbands is preferred. The transducermay be manufactured and tuned to exhibit a response characteristic asshown in FIG. 6, in which the lower hump 60 of the responsecharacteristic is centered about the transmitted fundamental frequencyf_(t), and the upper hump 62 is centered about the received harmonicfrequency f_(r) of the response passband. The transducer responsecharacteristic of FIG. 7 is preferred, however, as the single dominantcharacteristic 64 allows the probe to be suitable for both harmonicimaging and conventional broadband imaging. The characteristic 64encompasses the transmitted fundamental frequency f_(t), and also theharmonic receive passband bounded between frequencies f_(L) and f_(c),and centered about frequency f_(r). As discussed above, a lowfundamental transmit frequency of 1.67 MHz will result in harmonicreturning echo signals at a frequency of 3.34 MHz. A responsecharacteristic 64 of approximately 2 MHz would be suitable for thesefundamental and harmonic frequencies.

Tissue and cells in the body alter the transmitted fundamental frequencysignals during propagation and the returned echoes contain harmoniccomponents of the originally transmitted fundamental frequency. In FIG.1 these echoes are received by the transducer array 112, coupled throughthe T/R switch 114 and digitized by analog to digital converters 115.The sampling frequency f_(s) of the A/D converters 115 is controlled bythe central controller. The desired sampling rate dictated by samplingtheory is at least twice the highest frequency f_(c) of the receivedpassband and, for the preceding exemplary frequencies, might be on theorder of at least 8 MHz. Sampling rates higher than the minimumrequirement are also desirable.

The echo signal samples from the individual transducer elements aredelayed and summed by a beamformer 116 to form coherent echo signals.The digital coherent echo signals are then filtered by a digital filter118. In this embodiment, the transmit frequency f_(t) is not tied to thereceiver, and hence the receiver is free to receive a band offrequencies which is different from the transmitted band. The digitalfilter 118 bandpass filters the signals in the passband bounded byfrequencies f_(L) and f_(c) in FIG. 7, and can also shift the frequencyband to a lower or baseband frequency range. The digital filter could bea filter with a 1 MHz passband and a center frequency of 3.34 MHz in theabove example. A preferred digital filter is a series of multipliers70-73 and accumulators 80-83 as shown in FIG. 9. This arrangement iscontrolled by the central controller 120, which provides multiplierweights and decimation control which control the characteristics of thedigital filter. Preferably the arrangement is controlled to operate as afinite impulse response (FIR) filter, and performs both filtering anddecimation. For example, only the first stage output 1 could becontrolled to operate as a four tap FIR filter with a 4:1 decimationrate. Temporally discrete echo samples S are applied to the multiplier70 of the first stage. As the samples S are applied, they are multipliedby weights provided by the central controller 120. Each of theseproducts is stored in the accumulator 80 until four such products havebeen accumulated (added). An output signal is then produced at the firststage output 1. The output signal has been filtered by a four tap FIRfilter since the accumulated total comprises four weighted samples.Since the time of four samples is required to accumulate the outputsignal, a 4:1 decimation rate is achieved. One output signal is producedfor every four input samples. The accumulator is cleared and the processrepeats. It is seen that the higher the decimation rate (the longer theinterval between output signals), the greater can be the effective tapnumber of the filter.

Alternatively, temporally separate samples are delayed by delay elementsτ and applied to the four multipliers 70-73, multiplied, and accumulatedin the accumulators 80-83. After each accumulator has accumulated twoproducts, the four output signals are combined as a single outputsignal. This means that the filter is operating as an eight tap filterwith a 2:1 decimation rate. With no decimation, the arrangement can beoperated as a four tap FIR filter. The filter can also be operated byapplying echo signals to all multipliers simultaneously and selectivelytime sequencing the weighting coefficients. A whole range of filtercharacteristics are possible through programming of the weighting anddecimation rates of the filter, under control of the central controller.The use of a digital filter provides the advantage of being quickly andeasily changed to provide a different filter characteristic. A digitalfilter can be programmed to pass received fundamental frequencies at onemoment, and harmonic frequencies at the next. The digital filter canthus be operated to alternately produce images or lines of fundamentaland harmonic digital signals, or lines of different alternatingharmonics in a time-interleaved sequence simply by changing the filtercoefficients during signal processing.

Returning to FIG. 1, to image just a non-fundamental frequency, thedigital filter 118 is controlled by the central controller 120 to passecho signals at a harmonic frequency for processing, to the exclusion ofthe fundamental frequency. The harmonic echo signals from the tissue aredetected and processed by either a B mode processor 37 or a contrastsignal detector 128 for display as a two dimensional ultrasonic image onthe display 50.

The filtered echo signals from the digital filter 118 are also coupledto a Doppler processor 130 for conventional Doppler processing toproduce velocity and power Doppler signals. The outputs of theseprocessors are coupled to a 3D image rendering processor 162 for therendering of three dimensional images, which are stored in a 3D imagememory 164. Three dimensional rendering may be performed as described inU.S. Pat. No. 5,720,291 and in U.S. Pat. Nos. 5,474,073 and 5,485,842,the latter two patents illustrating three dimensional power Dopplerultrasonic imaging techniques. The signals from the contrast signaldetector 128, the processors 37 and 130, and the three dimensional imagesignals are coupled to a video processor 140 where they may be selectedfor two or three dimensional display on an image display 50 as dictatedby user selection.

It has been found that harmonic imaging of tissue and blood can reducenear field clutter in the ultrasonic image. It is believed that theharmonic response effect in tissue is dependent upon the energy level ofthe transmitted waves. Near to an array transducer which is focused at agreater depth, transmitted wave components are unfocused and ofinsufficient energy to stimulate a detectable harmonic response in thenear field tissue. But as the transmitted wave continues to penetratethe body, the higher intensity energy will give rise to the harmoniceffect as the wave components begin to focus. While both near and farfield regions will return a fundamental frequency response, clutter fromthese signals is eliminated by the passband of the digital filter 118,which is set to the harmonic frequency band. The harmonic response fromthe tissue is then detected and displayed, while the clutter from thenear field fundamental response is eliminated from the displayed image.

FIGS. 2, 3, 4, and 5 illustrate some of the properties of harmonicreturn signals which can be utilized to advantage in ultrasonic imaging.It should be appreciated that several of these properties and theirinteractions are not yet fully and commonly understood among thescientific community, and are still the subject of research anddiscussion. FIG. 2 illustrates the spatial response, and specificallythe main lobe and sidelobes, of fundamental and harmonic signalsreceived by a transducer array 112. In this illustration the array isdirected to image an area of the body behind the ribs, such as theheart, and the main lobe is seen to extend between ribs 10 and 10'.Overlying the ribs is a tissue interface 12, as from a layer of fatbetween the skin and ribs. The FIGURE shows a main lobe of thefundamental signals FL1, and on either side of the main lobe aresidelobes FL2 and FL3. The FIGURE also shows the main lobe HL1 of aharmonic of the fundamental frequency, and sidelobes HL2 and HL3 of theharmonic main lobe.

In this example it is seen that the main lobe of the fundamental echoesis wide enough to encompass portions of the ribs 10,10'. Accordingly,acoustic energy at the fundamental can be reflected back toward thetransducer 112 as indicated by the arrow 9. While some of the energy ofthis reflection may travel back to and be received directly by thetransducer, in this example some of the reflected energy is reflected asecond time by the tissue interface 12, as indicated by arrow 9'. Thissecond reflection of energy reaches the other rib 10', where it isreflected a second time as shown by arrow 9" and travels back to and isreceived by the transducer 112.

Since the intent of this imaging procedure is to image the heart behindthe ribs, these echoes reflected by the ribs are unwanted artifactswhich contaminate the ultrasonic image. Unwanted echoes which arereflected a number of times before reaching the transducer, such asthose following the paths of arrows 9,9',9", are referred to asmultipath artifacts. Together, these artifacts are referred to as image"clutter", which clouds the near field and in some cases all of theimage. This near field haze or clutter can obscure structure which maybe of interest near the transducer. Moreover, the multipath artifactscan be reproduced in the image at greater depths due to the lengthymultiple paths traveled by these artifacts, and can clutter and obscureregions of interest at greater depths of field.

But when only the harmonic return signals are used to produce theultrasonic image, this clutter from the fundamental frequencies isfiltered out and eliminated. The main lobe HL1 of the received harmonicechoes is narrower than that of the fundamental, and in this examplepasses between the ribs 10,10' without intersecting them. There are noharmonic returns from the ribs, and no multipath artifacts from theribs. Thus, the harmonic image will be distinctly less cluttered andhazy than the fundamental image, particularly in the near field in thisexample.

FIG. 3 shows a second example in which the main lobes of both thefundamental and harmonic returns do not intersect the ribs, and theproblem discussed in FIG. 2 does not arise. But in this example the ribs10, 10' are closer to the skin surface and the transducer 112. While themain lobes do not intersect the ribs, the sidelobes FL2 of thefundamental do reach the ribs, allowing sidelobe energy to be reflectedback to the transducer as shown by reflection path 9. Again, this willproduce clutter in the fundamental image. But the smaller and narrowersidelobes HL2 of the received harmonic energy do not reach the ribs.Again, the harmonic image will exhibit reduced clutter as compared tothe fundamental image.

FIG. 4 illustrates the fundamental and harmonic beam patterns in aperspective which is across the lobes of FIGS. 2 and 3, that is, acrossthe axis of the transducer. This drawing illustrates the relativeamplitude responses of the fundamental and second harmonic beampatterns. Illustrated are the dynamic response DRF between the main(FL1) and first sidelobe (FL2) of the fundamental component of the soundbeam, and the dynamic response DRH between the main (HL1) and firstsidelobe (HL2) of the second harmonic component. If responses due to themain lobes are considered desired signal responses, and responses due tothe sidelobes are considered to be clutter or noise, the signal to noiseratio of the harmonic is greater than that of the fundamental. That is,there is relatively less sidelobe clutter in a harmonic image than inthe corresponding fundamental image of the same transmission, orDRH>DRF.

FIG. 5 illustrates another comparison of the properties of fundamentaland harmonic signals, which is the relative amount of energy (in unitsof acoustic pressure P) emanating from increasing depths Z in the bodyat the fundamental and second harmonic frequencies. The curve denotedFund. shows the buildup of propagated acoustic energy at the fundamentalfrequency. While the curve is seen to peak at the focus of the arraytransducer, it is seen that there is nonetheless an appreciable amountof fundamental energy at the shallower depths before the focal region.In comparison, there is comparatively much less energy, and a lesserbuildup of energy, at the harmonic frequency propagated at these lesserdepths of field. Hence, with less energy available for multipathreverberation and other aberrations, there is less near field clutterwith harmonic imaging than with imaging the fundamental echo returnsfrom the same transmission.

FIG. 8 illustrates the bands of received signals and the digital filterof a typical FIG. 1 embodiment of the present invention for atransmitted signal of four cycles of a 1.67 MHz acoustic wave.Transmitting multiple cycles narrows the bandwidth of the transmittedsignal; the greater the number of cycles, the narrower the bandwidth. Inresponse to this transmission, the transducer 112 receives a fundamentalsignal in a bandwidth 90, which is seen to peak at the transmittedfrequency of 1.67 MHz. As the fundamental frequency band rolls off, theharmonic band 92 comes up, and is seen to exhibit a peak return at theharmonic frequency of 3.34 MHz. The received signals are applied to adigital filter with a passband characteristic 94, which is seen to becentered around the harmonic frequency of 3.34 MHz. As FIG. 8 shows,this passband will substantially suppress signals at the fundamentalfrequency while passing the harmonic signals on to further processingand image formation. When imaging the heart in this manner, it has beenfound that the harmonic response of the endocardial tissue of the heartis quite substantial, and harmonic tissue images of the heart show aclearly defined endocardial border.

Other signal processing techniques besides filtering may be used toseparate out harmonic signals from received echo information such ascancellation of the fundamental frequencies in a broadband signal,leaving only the harmonic frequencies. For example, U.S. Pat. No.5,706,819 discloses a two pulse technique, whereby each scanline isinsonified by consecutive fundamental frequency pulses of opposite phasein rapid succession. When the resultant echoes are received from the twopulses and combined on a spatial basis, the fundamental frequencies willcancel and the nonlinear or harmonic frequencies will remain. Thus, theharmonic frequencies are separated from the broadband echo signalswithout the need for a filter circuit.

FIG. 10 shows a portion of a preferred embodiment of the presentinvention in block diagram form, from the beamformer output through tothe image display. This embodiment not only produces harmonic images oftissue and blood flow, but also overcomes signal dropout deficiencies ofconventional imaging systems which arise when imaging patients withdifficult to image pathology. Additionally, this embodiment reduces anartifact of coherent ultrasound images known as speckle. In FIG. 10, thesignal and data lines connecting the blocks of the block diagram allrepresent multi-conductor digital data paths, as the processor of theillustrated embodiment is entirely digital. Scanline echo data from thebeamformer 116 is applied in parallel to the two channels 30a,30b of theprocessor illustrated in FIG. 10, one of which is a high frequencychannel and the other of which is a low frequency channel. Each channelof the processor has a normalization stage 32,132 which multiplies thescanline data by a scale factor on a sample by sample basis to producegain or attenuation that can vary with the depth of the body from whicheach sample returned. The scale factor for each channel is provided bynormalization coefficients stored in or generated by coefficientcircuits 32,132, which in a preferred embodiment are digital memories.As the multiplying coefficients are changed along the sequence ofscanline echoes, depth dependent gain or attenuation is produced.

The function of the normalization stages is two-fold. One is tocompensate for a transducer aperture which expands with depth of scan.As signals from an increasing number of transducers are used withincreasing depth, the magnitude of the summed beamformed signals willincrease. This increase is offset by reduced gain (increasedattenuation) in the normalization stage, in proportion to the rate atwhich channels are added to the beamforming process, so that theresultant echo sequence will be unaffected by the changing aperture.

The second function of the normalization stages is to equalize thenominal signal amplitudes of the two channels 30a,30b. The nominalsignal amplitudes of the passbands of the two channels are desirably thesame, so that the original relative signal levels will be preservedafter the passbands are summed to create the full harmonic passband. Butultrasound signals are subject to depth dependent attenuation whichvaries with frequency, higher frequencies being more greatly attenuatedwith depth than lower frequencies. To account for this depth dependentattenuation the coefficients for the normalization stages provide signalgain which increases with depth. Since the two channels employ differentfrequency passbands, the depth dependent gain of the two channelsdiffers from one channel to the other. In particular, the rate of gainincrease for the higher frequency passband channel is greater than thatof the lower frequency passband channel. This is illustrated in FIG. 11,which, for purposes of illustration, shows the normalization gaincharacteristic of the higher frequency passband channel separated intotwo components. The depth dependent characteristic 200 offsets theeffect of an increasing aperture in the channel, and the depth dependentcharacteristic 202 compensates for depth dependent signal attenuation.The low frequency passband channel may also have a depth dependent gaincharacteristic but with a different characteristic 202 for the differentrate of attenuation of the lower frequencies. The high frequencypassband channel has a similar but more rapidly increasing depthdependent gain characteristic to account for the more rapid rate ofattenuation of the higher frequencies. Each depth dependent gaincharacteristic 202 is chosen to offset the effect of depth dependentgain for the particular frequency passband employed by that channel.

In a preferred embodiment the coefficients of the coefficient circuitsapply a gain or attenuation characteristic which is a combination of thetwo characteristics 200,202. Preferably, the coefficient memories 32,132store multiple combined gain curves which are changed with memoryaddressing to match scanhead characteristics or the type of signalsbeing processed (2D or Doppler). The rate of gain change may becontrolled by the rate at which the coefficients are changed for themultiplier of each normalization stage 30,130.

The normalized echo signals in each channel are coupled to quadraturebandpass filters (QBPs) in each channel. The quadrature bandpass filtersprovide three functions: band limiting the RF scanline data, producingin-phase and quadrature pairs of scanline data, and decimating thedigital sample rate. Each QBP comprises two separate filters, oneproducing in-phase samples (I) and the other producing quadraturesamples (Q), with each filter being formed by a plurality ofmultiplier-accumulators (MACs) implementing an FIR filter. One such MACis shown in FIG. 12. As an echo sample of the scanline data is appliedto one input of a digital multiplier 210 a coefficient is applied to theother multiplier input. The product of the echo sample and the weightingcoefficient is stored in an accumulator 212 where it may be accumulatedwith previous products. Other MACs receive the echo samples at differentphases and likewise accumulate weighted echo samples. The accumulatedoutputs of several MACs can be combined, and the final accumulatedproduct comprises filtered echo data. The rate at which accumulatedoutputs are taken sets the decimation rate of the filter. The length ofthe filter is a product of the decimation rate and the number of MACsused to form the filter, which determine the number of incoming echosamples used to produce the accumulated output signal. The filtercharacteristic is determined by the values of the multiplyingcoefficients. Different sets of coefficients for different filterfunctions are stored in coefficient memories 38,138, which are coupledto apply selected coefficients to the multipliers of the MACs. The MACseffectively convolve the received echo signals with sine and cosinerepresentative coefficients, producing output samples which are in aquadrature relationship.

The coefficients for the MACs which form the I filter implement a sinefunction, while the coefficients for the Q filter implement a cosinefunction. For bandpass filtering, the coefficients of the active QBPsadditionally implement a low pass filter function that is frequencyshifted to form, in combination with the sine (for I) and cosine (for Q)functions, a bandpass filter for the quadrature samples. In the instantexample, QBP₁ in channel 30a is producing I and Q samples of thescanline data in a first, low frequency passband, and QBP₂ in channel30b is producing I and Q samples of the scanline data in a second,higher frequency passband. Thus, the spectrum of the original broadbandecho signals is divided into a high frequency band and a low frequencyband. To complete the dropout and speckle reduction process, the echodata in the passband produced by QBP₁ of channel 30a is detected by adetector 40₁ and the detected signals are coupled to one input of asummer 48. In a preferred embodiment detection is performed digitally byimplementing the algorithm (I² +Q²)^(1/2). The echo data in thecomplementary passband produced by QBP₂ of channel 30b is detected by adetector 40₂ and these detected signals are coupled to a second input ofthe summer 48. When the signals of the two passbands are combined by thesummer 48, the decorrelated signal dropout and speckle effects of thetwo passbands will at least partially cancel, reducing the signaldropout and speckle artifacts in the 2D image created from the signals.

Following the detector in each subchannel is a gain stage formed bymultipliers 44₁,44₂ which receive weighting coefficients fromcoefficient memories 42₁,42₂. The purpose of this gain stage is topartition the balance of analog and digital gains in the ultrasoundsystem for optimal system performance. Some of the gains in the echosignal path may be automatically implemented by the ultrasound system,while others, such as manual gain control and TGC gain, may becontrolled by the user. The system partitions these gains so that theanalog gains preceding the ADCs (analog to digital converters) of thebeamformer are adjusted optimally for the dynamic input range of theADCs. The digital gain is adjusted to optimize the brightness of theimage. The two gains together implement gain control changes effected bythe user.

In the preferred embodiment the gain imparted to the scanline signals bythe multipliers 44₁,44₂ is selected in concert with the gain of thepreceding normalization stage 34,134 in the channel. The gain of eachnormalization stage is chosen to prevent the attainment of saturationlevels in the QBPs, as may occur when strong signals from contrastagents or harmonic imaging are being received. To prevent saturationlevels the maximum gain of the normalization stage is controlled, andany reduction imposed by reason of this control is restored by the gainof the succeeding multiplier 44₁,44₂.

The gain function provided by these multipliers could be performedanywhere along the digital signal processing path. It could beimplemented by changing the slope of the compression curves discussedbelow. It could also, for instance, be performed in conjunction with thegains applied by the normalization stages. This latter implementation,however, would eliminate the ability to effect the saturation controldiscussed above. The present inventors have found implementation of thisgain function to be eased when provided after detection, and in thepreferred embodiment by use of a multiplier after detection.

The signals produced by the gain stages 44₁,44₂ generally exhibit agreater dynamic range than may be accommodated by the display 50.Consequently, the scanline signals of the multipliers are compressed toa suitable dynamic range by lookup tables. Generally the compression islogarithmic, as indicated by log compression processors 46₁,46₂. Theoutput of each lookup table is proportional to the log of the signalinput value. These lookup table are programmable so as to provide theability to vary the compression curves, and the brightness and dynamicrange of the scanline signals sent on for display.

The present inventors have found that the use of log compression toscale the echo signals can affect low level signals near the baseline(black) level of the signal dynamic range by exacerbating the degree andthe number of echoes with components at the black level, a manifestationof the destructive interference arising from the speckle effect of thecoherent ultrasonic energy. When the echo signals are displayed, many ofthem will be at the black level, and appear in the image to have beenundetected or dropped out. The embodiment of FIG. 10 reduces thisproblem by producing separate, partially decorrelated versions of theecho signals in the two channels 30a,30b. This embodiment partiallydecorrelates the echo signal versions by separating the echo signalcomponents into two different passbands as shown in FIG. 13. The twopassbands can be completely separated or, as shown in this example,overlapping. In this example, the lower passband 300a is centered abouta frequency of 3.1 MHz, and the higher passband 300b is centered about afrequency of 3.3 MHz, a center frequency separation of only 200 kHz.Even this small degree of separation has been found sufficient todecorrelate the signal components of the two passbands sufficiently suchthat black level signal dropout in one passband will frequently notalign in frequency with its corresponding component in the otherpassband. Consequently, when these decorrelated replicas of the sameecho signal are combined by the summer 48, the signal dropout andspeckle artifacts will be markedly reduced. This is especiallysignificant when trying to image fine structures at deep depths in thebody, such as the endocardium. A harmonic image of the endocardium issignificantly improved by the artifact elimination effects of theembodiment of FIG. 10.

As discussed previously the signal gain of the two passbands 300a,300bof FIG. 13 can be matched to preserve the original signal levels aftersummation. However, in a preferred embodiment, the lower frequencypassband is processed with less dynamic range than the higher frequencypassband as shown in FIG. 13. This has the effect of suppressing thefundamental frequency contributions of the lower frequency passband(which contains more fundamental frequency components than the higherfrequency band.) This is accomplished as a component of differentcompression characteristics in the log compression processors 46₁,46₂,or elsewhere in the channels 30a,30b subsequent to the separation of thebroadband signal into separate passbands.

The processed echo signals at the output of the summer 48 are coupled toa lowpass filter 52. This lowpass filter, like the QBPs, is formed bycombinations of multiplier-accumulators with variable coefficients,arranged to implement an FIR filter, to control the filtercharacteristic. The lowpass filter provides two functions. One is toeliminate sampling frequency and other unwanted high frequencycomponents from the processed echo signals. A second function is tomatch the scanline data rate to the vertical line density of the display50, so as to prevent aliasing in the displayed image. The FIR filterperforms this function by selectively decimating or interpolating thescanline data. The filtered echo signals are then stored in an imagememory 54. If the scanlines have not yet been scan converted, that is,they have r,θ coordinates, the scanlines are scan converted torectilinear coordinates by a scan converter and greyscale mappingprocessor 56. If scan conversion has been performed earlier in theprocess, or is not needed for the image data, the processor 56 maysimply convert the echo data to the desired greyscale map by a lookuptable process. The image data may then be stored in a final image memoryor sent to a video display driver (not shown) for conversion to displaysignals suitable for driving the display 50.

It will be appreciated that, due to the advantage of the quickprogrammability of a digital filter, the processing described above canbe performed in an embodiment which utilizes a single one of thechannels 30a, 30b to process the echo data from a scanline twice toalternately produce a line of signals for each of the two passbands in atime-interleaved fashion. However, the use of two parallel channelsaffords twice the processing speed, enabling harmonic images to beproduced in real time and at twice the frame rate of a time multiplexedembodiment.

Harmonic images produced from high frequency signals can suffer fromdepth dependent attenuation as the echo signals return from increasingdepths in the body. Lower frequency fundamental signals may experienceless attenuation, and hence in some cases may exhibit better signal tonoise ratios at greater depths. The embodiment of FIG. 14 takesadvantage of this characteristic by blending fundamental and harmonicimage data in one image. It is possible, for instance, to create anormal tissue image of the heart from fundamental frequencies, andoverlay the fundamental frequency tissue image with a harmonic tissueimage of the heart to better define the endocardial border in thecomposite image. The two images, one from fundamental frequencycomponents and another from harmonic frequency components, may be formedby alternately switching the digital filter 118 between fundamental andharmonic frequencies to separately assemble fundamental and harmonicimages, or by employing the two parallel filters of FIG. 10 with twopassbands, one set to pass fundamental frequencies and the other set topass harmonic frequencies. In FIG. 14, the filter of channel 30a is setto pass fundamental signal frequencies, and echo signals passed by thischannel are stored in a fundamental image memory 182. Correspondingly,harmonic signal frequencies are passed by channel 30b and stored in aharmonic image memory. The fundamental and harmonic images are thenblended together by a proportionate combiner 190, under control of ablend control 192. The blend control 192 may automatically implement apre-programmed blending algorithm, or one directed by the user. Forexample, the proportionate combiner 190 may create a blended image whichuses only echo data from the harmonic image at shallow depths, thencombines echo data from both image at intermediate depths, and finallyonly uses echo data of the fundamental image at deep depths. Thiscombines the reduced clutter benefit of harmonic echo data at shallowdepths and the greater penetration and signal to noise ratio offundamental echoes received from deeper depths, while affording a smoothtransition from one type of data to the other at intermediate depths.Other combining algorithms are also possible, such as simply switchingfrom one type of data to another at a predetermined depth, or outlininga region of the image to be displayed with one type of data while theremainder of the image is displayed using the other type of data.

It is also possible to employ the two parallel filters and blend thecomponents together before image formation, thereby adding acontrollable component of the harmonic echo signals to the fundamentalfrequency signals to enhance the resultant image. Such an embodimentcould eliminate the need for separate fundamental and harmonic imagememories and would process the signal components directly to a blendedimage memory.

A third technique for producing blended images is to receive eachscanline of the image through a depth-dependent, time varying filter.Such filters are well known for improving the signal to noise ratio ofreceived echo signals in the presence of depth dependent attenuation asshown, for instance, in U.S. Pat. No. 4,016,750. For the production ofblended fundamental and harmonic images, the passband 210 of a timevarying filter is initially set to pass harmonic frequencies f_(h), asshown in FIG. 15, as echo signals begin to be received from shallowdepths. When it becomes desirable to begin supplementing the image withfundamental signal components at deeper depths, the passband 210undergoes a transition to lower frequencies, eventually moving to thefundamental frequencies f_(f) as shown by passband 212 in FIG. 15. Inthe case of a digital filter such as that shown in FIG. 9, the change inpassband frequencies is effected by changing the filter coefficientswith time. As the filter undergoes this transition, the passband passesfewer harmonic frequencies and greater fundamental frequencies untileventually, if desired, the passband is passing only fundamentalfrequencies at the maximum image depth. By receiving each scanlinethrough such a time varying filter, each line in the resultant image cancomprise harmonic frequencies in the near field (shallow depths),fundamental frequencies in the far field (deepest depths), and a blendof the two in between.

A fourth technique for producing blended images is to transmit andreceive twice along each scanline. One transmission is at a fundamentalfrequency and is followed by reception of echoes at a harmonicfrequency. Another transmission is at a fundamental frequency, followedby reception at the fundamental frequency. The two fundamentaltransmission frequencies may be the same or, if desired, may bedifferent fundamental frequencies. The harmonic and fundamental echoesare then combined along the scanline in the desired proportions to forma blended scanline, and an image field of such scanlines are produced toform a blended image.

Harmonic tissue images of moving tissue can also be formed by processingthe received harmonic tissue echo signals with the processor describedin U.S. Pat. No. 5,718,229, entitled MEDICAL ULTRASONIC POWER MOTIONIMAGING.

Thus, the present invention encompasses an ultrasonic imaging system forimaging the nonlinear response of tissue and fluids of the body toultrasound by transmitting a fundamental frequency signal, receiving anecho signal from the tissue at a non-fundamental, preferably harmonic,frequency, detecting the non-fundamental frequency echo signals, andforming an image of the tissue and fluids from the non-fundamentalfrequency echo signals. As used herein the term harmonic also refers toharmonic frequencies of higher order than the second harmonic and tosubharmonics, as the principles described herein are equally applicableto higher order and subharmonic frequencies.

What is claimed is:
 1. An ultrasonic diagnostic imaging system forimaging the harmonic response of tissue inside a body with reducedclutter, comprising:a transit controller operable to cause the elementsof an array transducer to transmit wave components which are ofinsufficient energy to stimulate a significant harmonic response in thenear field, and which wave components are focused to develop higherintensity energy at a greater depth; an array transducer responsive tosaid transmit controller for transmitting ultrasonic energy into thebody at a fundamental frequency and responsive to said transmittedultrasonic energy, for receiving ultrasonic echo signals from tissue ata harmonic of said fundamental frequency; a beamformer which processesecho signals from the transducer elements of said array transducer toform coherent echo signals: a circuit for passing harmonic frequencyecho signals from tissue to the substantial exclusion of signals at saidfundamental frequency; and an image processor, responsive to harmonicfrequency echo signals from tissue passed by said circuit, for producingan ultrasonic tissue harmonic image, whereby multipath clutter in saidtissue harmonic image is substantially reduced.
 2. The ultrasonicdiagnostic imaging system of claim 1, wherein said array transducer fortransmitting and for receiving comprises an ultrasonic transducer arrayprobe.
 3. The ultrasonic diagnostic imaging system of claim 2, whereinsaid ultrasonic transducer array probe comprises a plurality oftransducer elements for transmitting ultrasonic energy at a fundamentalfrequency and for receiving ultrasonic echo signals at a harmonic ofsaid fundamental frequency.
 4. The ultrasonic diagnostic imaging systemof claim 3, wherein said transducer elements exhibit a responsecharacteristic which encompasses both said fundamental frequency andsaid harmonic of said fundamental frequency.
 5. The ultrasonicdiagnostic imaging system of claim 1, wherein said circuit for passingultrasonic echo signals at a harmonic of said fundamental frequencycomprises a filter defining a passband which includes said harmonicfrequency to the substantial exclusion of said fundamental frequency,including the substantial exclusion of multipath clutter at saidfundamental frequency.
 6. The ultrasonic diagnostic imaging system ofclaim 5, wherein said filter comprises a programmable digital filter. 7.The ultrasonic diagnostic imaging system of claim 1, wherein said imageprocessor includes a B mode processor which produces tissue harmonicimages.
 8. The ultrasonic diagnostic imaging system of claim 7, whereinsaid B mode processor includes an amplitude detector for detecting theenvelope of said harmonic echo signals.
 9. The ultrasonic diagnosticimaging system of claim 1, wherein said tissue comprises naturallyoccurring tissue and cells of the body.
 10. A method for producing anultrasonic image from the harmonic response of tissue in the interior ofthe body comprising the steps of:transmitting ultrasonic energy from theelements of an array transducer into the body at a fundamental frequencyin wave components emanating over a plurality of elements in the nearfield and which wave components become focused to develop harmonicfrequency components at a greater death; receiving ultrasonic echosignals by said array transducer which have been returned by said tissueat a harmonic of said fundamental frequency; forming coherent echosignals from said received ultrasonic echo signals; passing tissueharmonic signals to the substantial exclusion of fundamental frequencysignals; and processing said tissue harmonic signals from whichfundamental frequency signals have been substantially excluded toproduce ultrasonic image display signals; and displaying said ultrasonicimage display signals, whereby multipath clutter in said tissue harmonicultrasonic image is substantially reduced.
 11. The method of claim 10,wherein the steps of transmitting and receiving comprise using anultrasonic probe with a transducer array to transmit fundamentalfrequency ultrasonic energy and receive tissue harmonic echo signals.12. The method of claim 11, wherein the step of using an ultrasonicprobe comprises the step of transmitting fundamental frequencyultrasonic energy and receiving tissue harmonic echo signals with thesame transducer elements.
 13. The method of claim 10, wherein the stepof passing tissue harmonic signals comprises filtering receivedultrasonic echo signals to pass signals at said harmonic of saidfundamental frequency to the substantial exclusion of said fundamentalfrequency, including the substantial exclusion of multipath clutter atsaid fundamental frequency.
 14. The method of claim 10, wherein saidprocessing step comprises B mode processing said tissue harmonicsignals.
 15. The method of claim 14, wherein said step of B modeprocessing includes the step of amplitude detecting said tissue harmonicsignals.
 16. An ultrasonic diagnostic imaging system for imaging theharmonic response of tissue inside a body which exhibits depth dependentattenuation of ultrasonic energy, comprising:a transmit controlleroperable to cause the elements of an array transducer to transmit wavecomponents, the energy of which is distributed over the array in thenear field and becomes focused to develop harmonic frequency componentsat a focal depth; a transducer array responsive to said transmitcontroller for transmitting ultrasonic energy into the body at afundamental frequency which is equal to or less than 5 MHz and which isresponsive to said transmitted ultrasonic energy for receivingultrasonic echo signals from tissue at a harmonic of said fundamentalfrequency which is equal to or less than 10 MHz; means for digitizingsaid received ultrasonic echo signals, a digital beamformer for formingcoherent echo signals from said digitized ultrasonic echo signals; afilter which passes, tissue harmonic echo signals to the substantialexclusion of fundamental frequency signals; and an image processor,responsive to said tissue harmonic echo signals from which fundamentalfrequency signals have been substantially excluded, for producing anultrasonic image from said tissue harmonic echo signals, whereby themultipath clutter of said image of tissue harmonic echo signals issubstantially reduced.
 17. The ultrasonic diagnostic imaging system ofclaim 16, wherein said tissue comprises tissue and cells of the body.18. The ultrasonic diagnostic imaging system of claim 16, wherein saidtransducer array transmits ultrasonic energy into the body at afundamental frequency which is equal to or less than 2.5 MHz; andwhereinsaid transducer array receives ultrasonic echo signals from tissue at aharmonic of said fundamental frequency which is equal to or less than 5MHz.
 19. The ultrasonic diagnostic imaging system of claim 16, whereinsaid transducer array transmits ultrasonic energy into the body at afundamental frequency which is less than 2 MHz; andwherein saidtransducer array receives ultrasonic echo signals from tissue at aharmonic of said fundamental frequency which is less than 4 MHz.
 20. Theultrasonic diagnostic imaging system of claim 16, wherein said filtercomprises a programmable digital filter which is programmed to pass aband of said tissue harmonic echo signals to the substantial exclusionof signals of said fundamental frequency, including the substantialexclusion of multipath clutter at said fundamental frequency.